Method to automatically tune MRI RF coils

ABSTRACT

A sweep generator 230 applies a range of frequencies to an rf coil 140 to detect the resonant frequency of a field generated by a magnet 125a,b. A frequency to current converter 220 applies an auxiliary magnetic field to tune an MRI apparatus to the resonant frequency of the rf coil. A flexible coil of one turn (300) or two or more turns (500) has a plurality of segments (301-307; 501-513). One of the belt has a contact k0, (k0&#39;), which is electrically connectable to one or more contacts k1, (k1&#39;), k2, (k2&#39;), etc. located between the ends of the segments. For each connection to successive contacts, the length of the coil and its inductance increases by the added impedance ΔL ij  between contacts k i  and k j . That increase of inductance is nullified by capacitors ΔC sij  located between segments.

BACKGROUND OF THE INVENTION

1. Field of the Invention

This invention relates in general to magnetic resonance imaging (MRI),and in particular to apparatus and methods for automatically tuning anMRI coil.

2. Description of Related Art

In an MRI process, a sample, such as a human being, is placed in a largemagnetic field (the B₀ field) that remains constant throughout the MRIprocess. The magnetic moment of nuclei in the body, in particular nucleiof hydrogen, become aligned with the magnetic field. Next the sample isexposed to an oscillating magnetic field having a selected frequency inthe radio frequency (rf) region of the electromagnetic spectrum, causingthe nuclei in the sample to resonate. The rf radiation is then switchedoff, but the nuclei continue to resonate resulting in the emission of rfradiation from the resonating nuclei. The emission is detected as an MRIsignal

The resonance frequency of the sample depends upon the strength of thelarge magnetic field. This frequency is called the Larmor frequency andis expressed by the relationship L=γH, where L is the Larmor frequency,H is the strength of the magnetic field, γ is a constant dependent uponthe particular nuclei. The oscillating rf field is generated by an rfcoil that encloses the sample. The frequency of the applied oscillatingrf field is chosen to be substantially the same as the Larmor frequency.

The rf coil may also be used to receive the resonating emissions fromthe sample. The inductance and capacitance of the rf coil determine thetuned frequency of the coil and the impedance of the coil. The coilimpedance is matched to the optimum source impedance of the preamplifierso that the noise figure of the preamplifier is minimized. The rf coilhas its maximum sensitivity for the detecting emitted rf radiation whenthe inductance and capacitance of the rf coil are chosen so that the rfcoil has a tuned frequency which is the same as the Larmor frequency ofemitted rf radiation.

Prior art systems have, under certain conditions, had problems tuningthe frequency of the rf coil to the Larmor frequency of the nuclei.Since the coil includes reactive elements (coil inductance andcapacitive elements) the impedance of the coil is frequency dependent.Coil impedance has real and imaginary components that vary withfrequency. Most coils have their impedance tuned and matched t omaximize the signal-to-noise ratio of the detected signal. Often it issufficient to adjust tuning and matching capacitors upon construction ofthe rf coil to match the source impedance of the preamplifier whichminimizes the preamplifier's noise figure.

For an rf receive coil with a fixed geometry, the signal-to-noise ratioof magnetic resonance signals from a sample increases approximatelylinearly with the magnetic field. The closer the rf receive coil is tothe sample, the larger the signal and the signal-to-noise ratio. Thus,for low fields it is very important that the receive coil be close tothe body. The greater the distance between the coil and body, the poorerthe MRI image.

The resonance frequency of a coil is determined by the reactive elementsof the coil. The subject inside the coil increases it resistance whichprimarily affects the coil's bandwidth. The closer the coil is to thebody, the larger the capacitive coupling to the body and the greater theshift in coil resonance frequency due to variances in the capacitivecoupling between the coil and the body. A coil having a variablegeometry will have a variable inductance. A chance in coil inductancedivided by the nominal coil inductance which is of the order of the Q ofthe loaded coil will severely detune the coil, resulting in reducedimage quality. The resonance frequency of the coil may vary greatly frompatient to patient due to changes in coil geometry from patient topatient.

Another factor affecting resonance frequency is variability in the valueof the main magnetic field. Superconducting magnets produce high, stablemagnetic fields (typically greater than 0.3T). Permanent magnets orconventional electromagnets can produce fields which often vary withtime due to temperature variations, resulting in changes in theresonance frequency. Drifts in the magnetic field for a permanent magnetcan be of the order of a thousand parts-per-million (PPM) per degreeCentigrade resulting in a shift in the Larmor frequency. Such driftsresult in the coil being severely detuned with respect to the Larmorfrequency of the MRI system, leading to poor signal-to-noise ratios and,hence, poor image quality.

There are a number of techniques for tuning coils to the resonancefrequency of the MRI system. See, for example, U.S. Pat. No. 4,897,604which shows an expandable rf coil composed of a two parts including amain section and a removable bridge segment. Different size bridgesegments change both the active and the physical circumference of thecoil. In U.S. Pat. No. 5,143,068 a flexible coil having a fixed physicalsize is tuned by an externally located coupling coil circuit that hasvariable capacitors C_(s) and C_(p). U.S. Pat. No. 4,791,372 also reliesupon variable capacitors C_(s) and C_(p).

Prior art systems have several limitations. For example, manuallyretuning an MRI coil for each patient is time-consuming and inefficient.The larger the frequency range over which tuning is required, thegreater the time required, the more the difficulty in achieving a match,and thus the greater the cost. Imaging must be delayed until tuning iscomplete. Such delays add an extra financial burden to health care costsince the time of the patient and the MRI operator are taken up with thetuning process.

In the current art a single coil is not well suited for use withdifferent sized patients because the frequency range for coil tuning andthe coil filling factor cannot be made optimal for all patients. Priorart attempts to overcome these limitations with a multiplicity ofdiscrete elements suffer from unreliability due to the number ofconnectors required between segments and are cumbersome to use since thecoil must be reconstructed for each patient. See U.S. Pat. No.4,897,604. Furthermore, each connection can add a series resistance tothe coil due to contact resistance, resulting in a degradation of thecoil signal to noise.

OBJECTS OF THE INVENTION

It is an object of this invention to provide an automatically tuned MRIsystem.

It is a further object to provide a single coil of variable activecircumference that is automatically tuned to the sample regardless ofthe circumference of the coil.

It is a further object to provide a coil of variable circumference thatcan be easily applied to patients having a variety of sizes.

It is a further object to provide a means of fine tuning the coil with aminimal addition of noise to the coil.

These and other objects of the invention will be understood by thoseskilled in the art with reference to the following summary and detaileddescription and the attached drawings.

SUMMARY OF THE INVENTION

In contrast to the prior art, the present invention sets the magneticfield in an Magnetic Resonance Imaging (MRI) system so that the Larmorfrequency of the nuclear spins is equal to the tuned frequency of the rfcoil. The primary magnetic field of the MRI system is generated by apermanent magnet, an electromagnet or a superconducting magnet. Anauxiliary magnetic coil generates an auxiliary magnetic field. Thecombined primary and auxiliary magnetic fields produce a net magneticfield such that the Larmor frequency of the nuclei being probed are atthe tuned frequency of the rf coil.

In a first embodiment of the present invention a sweep generatorproduces a variable rf frequency signal which is sent to the rf coil viaa signal splitting device such as a quadrature hybrid. When theinstantaneous frequency of the sweep generator is identical to the tunedfrequency of the rf coil, the amount of signal propagated into the rfcoil is maximized. Consequently, the amount of signal appearing at adetection port of the signal splitting device will be minimized. Asignal measuring means monitors the detection port of the signalsplitting device and sends the amplitude of the detected signal to aresponse frequency identifier means. The response frequency identifiermeans identifies the frequency of the signal from the sweep generator atthe instance of maximum energy transfer into the rf coil. This frequencyis the tuned frequency of the rf coil. The frequency measured by theresponse frequency identifier means is propagated to a frequencyconverter means which adjusts the strength of the auxiliary magneticfield. The gain of the frequency converter means is selected so that thenet magnetic field created by the primary and auxiliary magnets andexperienced by the nuclear spins results in the Larmor frequency of thespins being equal to the tuned frequency of the rf coil.

In a second embodiment of the present invention the Larmor frequency ofthe subject is determined by applying at least one wide-band rf pulse tothe subject and detecting the resulting rf emissions. The magnetic fieldcreated by the auxiliary magnetic field coils is then changed so thatthe nuclear spins experience a different net magnetic field.Commensurate with the change in the net magnetic field, the transmitterand receiver frequencies are changed by an amount which is directlyproportional to the change of the auxiliary magnetic field so that thedetected magnetic resonance signals retain their frequency offsets withrespect to the transmitter and receiver frequencies. The strength of thedetected magnetic resonance signal is measured at the new net field andis compared to the strength of the magnetic resonance signal prior tothe change in the auxiliary field. The entire process is repeated for aseries of different net magnetic fields to identify the magnetic fieldwhich gives the strongest detected magnetic resonance signal. Thestrongest detected magnetic resonance signal will be obtained when theLarmor frequency of the nuclear spins is equal to the tuned frequency ofthe rf coil.

Another embodiment of the invention uses a flexible coil formed as amulti-layer continuous belt. One layer of the coil is made ofelectrically insulating material. Over the insulating layer is a secondlayer that includes one or more rf coils for generating or receiving anrf MRI imaging signal. The belt has a minimum physical length and isadjustable in discrete serial active segments to a maximum physicallength. The number of connected serial segments determines the activelength of the belt. The active length can be less than or equal to themaximum physical length. The minimum length of the belt includes a firstactive segment that has an inductance and a capacitance. They establisha resonant frequency for the minimum length such that the net reactanceis zero. Electrical contacts on one end of the belt are connected tocorresponding contacts in the segments to provide means for adjustingthe length of the belt to an active length that is an intermediatephysical length compatible with the circumference of a sample. Eachintermediate length includes one or more active serial segments of thebelt that extend from the minimum length to the maximum length. Eachserial segment has an incremental inductor and an incremental capacitorthat substantially cancel one another so that the resonant frequency ofthe coil at each intermediate length and at the maximum length issubstantially the same as the resonant frequency for the minimum length.

The belt provides an automatic coarse tuning of the rf coil(s) fordifferent size samples. The invention provides further means for finetuning the rf coil(s) using external controls.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a general schematic view of an apparatus used to practice themethod of the invention.

FIG. 2 is a detailed schematic of the components of an MRI apparatusused to practice the invention.

FIG. 3a is a combined mechanical and electrical schematic of theflexible coil.

FIG. 3b is an equivalent electrical circuit of a flexible coil.

FIG. 3c is an end view of the flexible coil at its minimum activelength.

FIG. 4 is an equivalent electrical circuit of the flexible coil.

FIG. 5 is a combined mechanical and electrical schematic of the flexiblecoil with two turns.

FIG. 6a is a switched capacitor tuning circuit for the flexible coil ofFIG. 3 and FIG. 5.

FIG. 6b is an equivalent electrical circuit of the flexible coil of FIG.3 and FIG. 5 with a phase tuning circuit.

DETAILED DESCRIPTION OF THE INVENTION 1. Tuning Method

Turning to FIG. 1, a subject 100 is placed on a support table 110 andpositioned in a homogeneous primary magnetic field generated by primarymagnets 120a and 120b and an auxiliary magnetic field generated byauxiliary electromagnets 125a, 125b. In the present embodiment primarymagnets 120a, 120b are permanent magnets that establish a field acrossthe subject 100. Other embodiments in which resistive or superconductingmagnets are employed are also possible. Auxiliary magnets 125a, 125bgenerate the auxiliary magnetic field parallel to the primary magneticfield and across the subject 100. A set of magnetic field gradient coils130 surround subject 100 and create magnetic field gradients ofpredetermined strength at predetermined times according to predeterminedMR pulse sequences. Gradient coils 130 generate magnetic field gradientsin three mutually orthogonal directions. At least one externalradio-frequency (rf) coil 140 (only one is shown in FIG. 1) alsosurrounds the region of interest of subject 100. In FIG. 1, gradientcoils 130 and rf coil 140 have a cylindrical shape with a diametersufficient to encompass the entire subject. Other geometries, such assmaller cylinders specifically designed for imaging the head or anextremity, can be used in alternative embodiments. Non-cylindricalgradient and radio-frequency coils, such as surface coils, may also beused.

External rf coil 140 radiates radio-frequency energy into subject 100 atpredetermined times and with sufficient power at a predeterminedfrequency so as to nutate a population of nuclear magnetic spins,hereinafter referred to as "spins", of subject 100 in a fashion wellknown to those skilled in the art. External rf coil 140 nutates spins bycreating a rotating magnetic field which is perpendicular to the primarymagnetic field and rotates at the Larmor frequency. If desired, externalrf coil 140 can also act as a receiver and detect the MR responsesignals that are stimulated by nutation. The system is operated undercontrol of the operator 112 and in accordance with a set of instructionsstored in a computer contained within an imaging electronics subsystem115.

The nutation resonates the spins at the Larmor frequency. The Larmorfrequency for each spin is directly proportional to the strength of thenet magnetic field experienced by the spin. The net field strength isthe sum of the static magnetic field generated by primary magnets 120a,120b and auxiliary magnets 125a, 125b and the local field generated bymagnetic field gradient coil 130. For heuristic purposes, the effect ofthe field gradient coil 130 will be ignored since it remains inactiveduring tuning of the magnets 120a, 120b and 125a, 125b to the resonantfrequency of the rf coil 140. In operation, the gradient coil 130provides a linearly variable magnetic field that distinguishes thelocation of rf signals generated at different locations of the subject100.

With reference to FIG. 2, a sweep generator 230 produces a variable rffrequency signal which is selectively sent to the rf coil 140 via asignal splitter 250 such as a quadrature hybrid or directional coupler.A switch 232 selectively connects the rf coil 140 to the imagingelectronics system 115 or a splitter 250. The switch 232 may be a manualswitch or a switch operable by suitable controls in the imagingelectronics system 115. Sweep generators are well known in the art. Theycan be set to any one of a number of frequencies and/or willautomatically step an output signal of a fixed peak-to-peak valuethrough a range of frequencies to produce output rf signals of the sameamplitude at the different frequencies defined by the chosen range.Quadrature hybrids splitters are also well known to those skilled in theart. Signals sent into a quadrature hybrid splitter exit the splitterthrough one or more ports. In the present embodiment, if the impedanceof the rf coil 140 matches the impedance of the resistor on the oppositeport (shown here to be 50 Ohms) then all the signal exits the hybrid viathe ports attached to the rf coil 140 and the resistor. Note that theimpedance of the rf coil 140 is closest to 50 Ohms at the tunedfrequency. Consequently, when the instantaneous frequency of the sweepgenerator 230 is identical to the tuned frequency of the rf coil 140,the amount of signal propagated into the rf coil is maximized and theamount of signal appearing at the detection port 235 of the signalsplitter 250 is minimized. A signal measuring means 260 monitors signalsappearing at the detection port of the signal splitting device and sendsa signal representative of the amplitude of the detected signal to aresponse frequency identifier means 270. The response frequencyidentifier means 270 identifies the frequency of the signal from thesweep generator 230 at the instance of maximum energy transfer into therf coil 140. This frequency is the tuned frequency of the rf coil 140.The applied current may be amplified by current amplifier 210. Thefrequency measured by the response frequency identifier means 270 ispropagated to a frequency converter means 220 which adjusts the strengthof the auxiliary magnetic field by applying an appropriate current toauxiliary magnets 125a, 125b. The gain of the frequency converter means220 is selected so that the net magnetic field created by the primaryand auxiliary magnets and experienced by the nuclear spins results inthe Larmor frequency of the spins being equal to the tuned frequency ofthe rf coil 140. The applied current may be amplified by currentamplifier 210. The frequency measured by the response frequencyidentifier means 270 is also propagated to the imaging electronicssubsystem 115 where it is used to set the frequency of the MR system'stransmitter and receiver. Once the rf coil 140 is tuned, switch 232 isset to connect rf coil 140 to imaging electronics system 115. Thefrequency of the rf signal is the resonant frequency that is output bythe frequency identifier means 270. In a variation of the firstembodiment of the present invention an external probe 141 can be placednear the rf coil 140 to probe the tuning of rf coil 140, instead ofemploying a direct electrical connection between the rf coil 140 and thesweep generator 230.

In a second embodiment of the present invention the tuned frequency ofthe rf coil 140 is determined without the use of a sweep generator 230.In the second embodiment, the Larmor frequency of the subject 100 isdetermined by applying at least one wide-band rf pulse to the subject100 and detecting the resulting rf emissions with rf coil 140. Themagnetic field created by the auxiliary magnetic field coils 125a, 125is then changed so that the nuclear spins experience a different netmagnetic field. Commensurate with the change in the net magnetic field,the transmitter and receiver frequencies of the imaging electronicssubsystem 115 are changed by an amount which is directly proportional tothe change of the auxiliary magnetic field so that the detected magneticresonance signals retain their frequency offsets with respect to thetransmitter and receiver frequencies. The strength of the detectedmagnetic resonance signal is measured at the new net field and iscompared to the strength of the magnetic resonance signal prior to thechange in the auxiliary field. The entire process is repeated for aseries of different net magnetic fields to identify the magnetic fieldwhich gives the strongest detected magnetic resonance signal. Thestrongest detected magnetic resonance signal will be obtained when theLarmor frequency of the nuclear spins is equal to the tuned frequency ofthe rf coil.

Although ideally suited for single rf coils used for both transmit andreceive, the first and second embodiments of the present invention maybe practiced with multiple rf coils if desired. For example, in an MRIsystem which has separate transmit and receive rf coils, the presentinvention can be selectively applied to only the receive coil sinceproper matching and tuning of a large transmit coil is not critical toimage quality whereas matching and tuning a receive coil is critical toimage quality. Alternatively, the tuned frequency of the receive andtransmit coils can be independently determined and the net magneticfield adjusted for one coil while an automatic or manual adjustment ofthe tuned frequency of the other coil is made. It should be noted that amistuned transmit coil requires more rf power whereas a mistuned receivecoil degrades the system's signal-to-noise ratio. Also, to avoid orminimize coupling between separate transmit and receive coils, duringtransmit the receive coil should be dynamically disabled and duringreceive the transmit coil should be dynamically disabled.

2. Flexible Body Coil

The foregoing tuning method may be used to tune a flexible body coil 300shown in FIG. 3. The flexible coil 300 has a fixed physical length andan active circumference that is adjustable to accommodate persons ofdifferent sizes. Because the inside volume of the coil is almostentirely filled with the patient, the signal pickup from the body ismaximized while the pickup of signals external to the body is minimized.Thus the coil signal-to-noise ratio, SNR, approaches the intrinsic limitfor pickup from the body.

The active coil length is adjusted by making contact of one end of thecoil to other fixed locations along the coil. The coil 300 also has abuilt-in matching circuit that minimizes the level of external tuningrequired. While some external tuning may be necessary, the bulk of thetuning required by changes in coil circumference is accomplishedmechanically by placing capacitors between the different contact pointsalong the coil's circumference. The values of the capacitors are chosenso their reactance cancels the increase in inductive reactance of thecoil on going from one connection to the next. If the coil is tuned to aresonance for the minimum active length, then the coil willautomatically be tuned to the same resonance frequency for all of thefixed locations. Fine tuning can then be accomplished by making one ofthe series capacitors variable.

The flexible body coil 300 is designed to maximize coupling to a widerange of patient sizes, to minimize noise generated by the coil and tominimize the amount of external tuning required to match the coil outputimpedance to the optimum source impedance of the preamplifier connectedto the coil. Because 50 ohm coaxial cables are often used, appropriatereactances are often attached to the preamplifier input in order to makethe optimum source impedance of the preamplifier to be 50 ohm with azero degree phase. For all further discussions, this will be thepreferred condition.

The flexible coil 300 can be wrapped around the patient like a belt andits circumference is adjustable to accommodate different size persons. Adiagram of a flexible body coil 300 is shown in FIGS. 3a, 3b, and 3c.FIG. 3a is the disconnected coil; FIG. 3b is the equivalent circuit forthe coil of FIG. 3a; FIG. 3c is the configuration of the coil for theminimum active length. The coil 300 has a single turn copper wire strip310 on a fiberglass insulation sheet 320, and a foam liner 330. The coil300 could be made with multiple turns about the body. Two capacitors,C_(s0) and C_(p), are connected in series with the copper wire strip 310within what is defined as the minimum active length, as shown in FIG.3c. C_(s0) could be a single capacitor or composed of several capacitorswhich are connected either in parallel or in series and whose effectivecombined capacitance equals C_(s0). The coil 300 has a minimal lengthdefined by contact of k0 with k1. When k0 is contacted to k1, the coilhas an inductance of L_(C0), (equal to the sum of L_(C1), L_(C2) andL_(C3)), and the capacitors C_(p) and C_(s0) are in series with L_(C0).Segments 301, 302 and part of 303 define the minimal length. A contactpoint is located at one end of segment 301 (k0') at about the mid-pointof segments 303, (k1), 304, (k2), 305, (k3), and 306, (k4), and, at thefar end of segment 307 (k5). Each incremental length is defined byadjacent halves of serial segments. The first incremental length isdefined by the length from k1 to k2. That first incremental lengthincludes part of segment 304. A capacitor ΔC_(s12) is chosen to nullifythe added reactance from ΔL₁₂ from segments 303 and 304. By choosingcapacitance values ΔC_(sij) whose reactance effectively cancels out theincremental reactance from the incremental inductance ΔL_(ij), thereactance around the loop of the coil 300 is maintained at effectivelythe same value for all active lengths. Thus, once the coil is tuned forthe minimum active length, the coil will remain tuned for all activelengths of the coil, as seen by making an effective circuit for thecoil.

The coil 300 can be characterized by an effective series LCR circuit,which is diagrammed in FIG. 4. The effective circuit for the resonantbody coil is a series inductor, L_(c), resistor, R_(s), and capacitorsC_(s) and C_(p), all in series. L_(c) is the effective series inductanceof the coil, which for the minimum active length in FIG. 3b is the sumof L_(C1), L_(C2) and L_(C3). R_(s) includes the series resistance ofthe copper wire strip 310 and the transformed resistance of the subject100 in the coil 300. C_(s) is equal to C_(s0) in FIG. 3b. The rf signalsare measured across C_(p) and a series inductor, L_(blk). The impedance,Z_(in), across C_(p) and L_(blk) of FIG. 4 is given in Equation 1:##EQU1##

Where j=√-1; X_(L) (=ωL_(c)) and X_(blk) (=ωL_(blk)) are the reactancefrom the coil series inductance, Lc, and the "blocking" inductor,L_(blk) ; and X_(p) (=1/ωC_(p)) and X_(s) (=1/ωC_(s)) are the capacitivereactance of the parallel, C_(p), and series, C_(s), capacitors, and ωis the radial frequency. The coil resonance frequency, f₀ =ω₀ /2 π, iswhen the series reactance around the coil circumference is zero whichoccurs when the reactance from the coil inductance, L_(c), matches thenet series capacitive reactance from C_(p) and C_(s) as given byEquation 2a. At resonance, the imaginary part of the impedance aroundthe coil circumference (X_(L) -X_(p) -X_(s)) is zero and the real partof the impedance across C_(p) and L_(blk), ReZ_(in) (f₀), is a maximumas given by Equation 2b. By choosing L_(blk) such that X_(blk) =X_(p),the imaginary component of the impedance across C_(p) and L_(blk) iszero. ##EQU2##

Thus if equations 2a and 2b are satisfied for the minimum active lengthand for each intermediate active length, the added net reactance is zeroand Equations 2a and 2b remain satisfied and unchanged for all contactpoints.

While nullifying the incremental reactive changes for the differentactive coil lengths maintains the coil resonance frequency, Equation 2bshows that if the series resistance, R_(S), changes for differentpatient sizes, the impedance, ReZ_(in) (ω₀) will change. In a singleturn coil we built for a 0.2T scanner, operating at 8.65 MHz, forpersons ranging from 100 lbs and 230 lbs and with waist circumferencesbetween 80 cm and 120 cm, the values of R_(s) ranged from about 0.2Ω to1Ω. From Equation 2b, at 8.65 MHz, using a value of about 4.7Ω forX_(p), the real part of the impedance will therefore be between about23Ω and 118Ω for all values of R_(s) measured in this study. For a lownoise pre-amplifier whose optimum input impedance is designed to be realand 50Ω, very little additional noise is added for the real inputimpedances between about one-half and two times this value. Thus it isnot necessary to make C_(p) variable in order to approximate the optimumSNR for different sized patients.

The flexible coil may include more than one turn. FIG. 5 shows a doubleturn solenoid flexible body coil 500. The coil 500 is adjustable withfour possible active circumferences (though there is no restrictionlimiting the number) with automatic coarse-tuning for all fourcircumferences. The coil 500 has thirteen segments, 501-513, andincludes a number of capacitors C_(A), C_(B), C_(C), C_(D), C_(E),C_(F), C_(G), C_(H), C_(I), C_(J), and C_(P) located between thesegments. Note that capacitor C_(J) may be a variable capacitor to allowfor some external tuning (but here it is shown as a fixed capacitor,C_(s0), in parallel with a variable capacitor, C_(sv)). Since there aretwo turns, two contacts are required for each segment, i.e. k0, k0', k1,k1', etc. The minimum active length is defined by contact of k0 and k0'with k1 and k1' respectively. The first intermediate active length isadded by contact of k0 and k0' to k2 and k2' respectively. As with thesingle turn coil, the mechanical tuning is achieved by making the netchange in the coil's reactance substantially zero when moving thecontacts of k0 and k0' from k1 and k1' to k2 and k2'. Furthermore, theelectrically insulating tape 530 separates the conductive sheets fromone another where they overlap. The copper plate 520 is connected toeach contact point k0, k0', k1, k1', etc. for added mechanical strength.

Ideally, once a coil has been tuned for a single patient for the minimumactive length of the coil, when the net reactance between connectionshas been tuned substantially to zero, no further tuning would berequired for increases in the active coil length. In practice,variability in the coil's resonance frequency associated with changes inthe coil's inductance due to slight geometrical differences betweenpatients and to slight changes in the mutual inductance coupling betweenthe coil and the magnet system results in a variability in the coil'sresonance frequency from patient to patient. For this reason, finetuning may be required for clinical applications. We have devised threeexternally controlled tuning methods. One tuning method involves tuningthe main magnetic field as described in the previous section on Tuning.The second tuning method involves externally fine tuning the reactanceof the active length of the coil to adjust the resonance frequency ofthe coil. The third tuning method involves externally tuning thereactive component of the coil's impedance to match the phase of thesource impedance into the preamplifier to zero. In general the impedanceof the flexible coil and any coaxial cable and connectors, as seen froma preamplifier, is ideally the same as the preamplifier source impedancebut may be in a range of about one half to twice the source impedance.

Externally tuning the reactance of the active component of the coil toadjust the resonance frequency of the coil can be accomplished byplacing a variable reactance in parallel or in series with capacitorC_(s0). One such variable reactive element is a variable capacitor,C_(sv). To maximize the signal-to-noise ratio of the coil, high Qvariable reactive elements are required. Two examples of high Q voltagecontrolled variable capacitors are: ultra-abrupt diode capacitors as acontinuously variable capacitor or switched capacitors using PIN diodesfor digital capacitance variability. One possible placement of theswitched capacitor unit is to connect the switched capacitors inparallel with capacitor C_(s0) in FIG. 4. In FIG. 5. the switchedcapacitors are represented by C_(sv).

FIG. 6a shows an example the switched capacitor tuning circuit 600containing two switched capacitors used to tune the flexible body coil300 or coil 500. More switched capacitors may be used to achieve a widerrange of tuning, increased resolution, or both. The capacitors 601 and602 are connected in series and the capacitors 603 and 604 are connectedin series. The capacitors 601-602 and 603-604 of the switched capacitorunit are connected in parallel with another capacitor 605, which couldalso be C_(s0) of FIG. 4 and FIG. 5. Diodes 610 and 612 are respectivelyconnected in parallel with capacitors 602 and 604. The diodes 610, 612are controlled by diode control circuit 620. The diode control circuitcontains reactive elements, such as inductors and capacitors, toelectrically isolate the coil from the external world at the operatingfrequencies of the coil. Each diode 610, 612 is either forward biased orreverse biased in accordance with a control voltage applied to the diodeby control circuit 620. So, the diodes 610, 612 switch capacitors 602,604, respectively, either into or out of the tuning circuit 600. Withthe two diodes 610, 612 the resonance frequency of the coil can be tunedto four possible values with the bias of the diodes 610, 612 being:forward, forward; forward, reverse; reverse, forward; reverse, reverse.PIN diodes are an excellent choice of switches because of their lowforward biased resistance, which currently can be as low about 0.25Ω at8.65 MHz. For fine tuning of the series capacitance of the coil's activelength, the capacitance values of capacitors 601-604 are much smallerthan that of capacitor 605. Furthermore, in order to minimize additionalseries resistance of the coil when the PIN diodes are forward biased,the reactance of diodes 601 and 603 should be much larger than theseries resistance of the forward biased PIN diodes. Making thecapacitance value of capacitor 605 much larger than the values ofcapacitors 601-604 also helps to reduce the addition to the totaleffective series resistance of the coil, R_(s), due to the presence inthe circuit of the diodes 610, 612.

Phase tuning can be performed between the coil output and thepreamplifier. FIG. 6b shows a phase tuning circuit placed immediatelyafter the coil. The phase seen by the pre-amplifier can be brought tozero by the addition of the appropriate reactive elements placed inseries with the coil's output, transforming the imaginary component ofthe coil's output impedance to zero with little affect on the realcomponent of the impedance. While phase tuning can be done with either avariable inductor or a variable capacitor, FIG. 6b depicts the variableelement to be a capacitor. The phase tuning elements shown in FIG. 6bare composed of a fixed inductor, L_(v), and a variable capacitor,C_(v). The variable capacitor is connected to a capacitance controllingdevice 630. For an ultra-abrupt diode variable capacitor, thecapacitance would be a voltage controlling device with appropriateelectronics. For example, for the coil tuned at resonance to 100Ω phasetuning over the range of ±45 degrees, the reactance of the inductor andcapacitor must respectively be equal to 50Ω and 100Ω which at 8.64 MHzcorrespond to an inductance, L_(v), of 0.92 μh and a capacitance, C_(v)of 184 pf. Hyper abrupt diodes are an excellent, high Q choice for avariable capacitor, C_(v), since they can be tuned over a large fractionof their nominal values and have low series resistances. Switchcapacitors could also be used to digitally vary C_(v).

While several presently preferred embodiments of the novel automatictuning method have been described in detail herein, many modificationsand variations will now become apparent to those skilled in the art. Itis, therefore, to be understood that the appended claims are intended tocover all such modifications and variations as fall within the truespirit of the invention.

What we claim is:
 1. In a magnetic resonance imaging apparatus having aprimary coil for generating a primary magnetic field in a firstdirection, an auxiliary coil for generating an auxiliary magnetic fieldin a direction parallel to the primary magnetic field and resulting in anet magnetic field in a first direction and an rf coil for transmittingand receiving an rf signal that generates an imaging signal in adirection substantially perpendicular to the direction of the netmagnetic field, a method for tuning the apparatus to the resonantfrequency of the rf coil comprising the steps of:placing a sample withinthe primary coil, auxiliary coil and rf coil of the magnetic resonanceimaging apparatus and causing the nuclear spins of the sample toresonate at a Larmor frequency proportional to the net magnetic field;applying radio frequency signals of one or more frequencies to the rfcoil; monitoring the transfer of energy into the rf coil in response tothe applied radio frequency signals; identifying the frequency at whichenergy transfer into the rf coil is maximized; varying the magnitude ofthe current applied to the auxiliary coil to change the net magneticfield so that the Larmor frequency of the sample is substantially thesame as the identified frequency at which energy transfer into the rfcoil is maximized.
 2. The method of claim 1 wherein the frequency of theradio signals applied to the rf coil are swept over a range of radiofrequencies to identify the resonance frequency.
 3. In a magneticresonance imaging apparatus having a primary coil for generating aprimary magnetic field in a first direction, an auxiliary coil forgenerating an auxiliary magnetic field in a direction parallel to theprimary magnetic field and resulting in a net magnetic field in a firstdirection, an rf coil for transmitting and receiving an rf signal thatgenerates an imaging signal in a direction substantially perpendicularto the direction of the net magnetic field, a transmitter for sending rfpulses to the rf coil, and a receiver to detect magnetic resonanceresponse signals from the rf coil, a method for tuning the apparatus tothe resonant frequency of the rf coil comprising the steps of:a) placinga sample within the primary coil, auxiliary coil and rf coil of themagnetic resonance imaging apparatus and causing the nuclear spins ofthe sample to resonate at a Larmor frequency proportional to the netmagnetic field; b) selecting a transmitter and receiver frequency whichhas a selected offset from the Larmor frequency of the sample; c)applying at least one rf pulse at the selected transmitter frequency tothe rf coil to nutate nuclear spins within the sample; d) detecting rfsignals in response to the applied rf pulse; e) measuring the amplitudeof the detected rf signal; f) varying the magnitude of the currentapplied to the auxiliary coil by a selected amount to change the netmagnetic field of the apparatus; g) varying the frequency of thereceiver and transmitter of the apparatus by an amount proportional tothe selected change in current applied to the auxiliary coil so that thedetected rf signal has an offset frequency substantially identical tothat selected in step b; h) repeating steps c-g a selected one or moretimes to determine the auxiliary coil current which gives the largestamplitude of the detected rf signal; i) setting the current in theauxiliary coils of the apparatus to the current found to give thelargest amplitude detected rf signal in step h.
 4. A tunable magneticresonance imaging apparatus comprising:a primary coil for generating astationary magnetic field in a first direction; an auxiliary coil forgenerating an auxiliary magnetic field in said first direction therebygenerating a net magnetic field in said first direction and establishinga resonant frequency in a sample in the net magnetic field that isproportional to the strength of a net magnetic field; an rf coil fortransmitting and receiving an rf signal that generates a second magneticfield in a second direction substantially perpendicular to the firstmagnetic field; signal splitting means for splitting a sweep generatorsignal between two output ports including one output port selectivelyconnectable to the rf coil; a sweep generator coupled to the signalsplitting means for generating rf signals of different frequencies; asignal measuring means coupled to the signal splitting means fordetecting the amount of rf signal which is propagated from the sweepgenerator through the signal splitting means to the rf coil and forgenerating an output signal proportional to the magnitude of thedetected rf signal; a response frequency identifier means coupled to thesweep generator and to the signal measuring means for the identifyingthe rf signal frequency corresponding to the maximum rf signalpropagated from the sweep generator to the rf coil; a frequencyconverter coupled to the response frequency identifier means forgenerating a current signal proportional to the frequency correspondingto the maximum rf signal propagated from the sweep generator to the rfcoil; a means for connecting the frequency converter to the auxiliarycoil to change the current of the auxiliary coil and thereby change theauxiliary magnetic field generated by the auxiliary coil.
 5. Theapparatus of claim 4 further comprising an imaging electronics systemcoupled to the output of the frequency identifier means for receivingmagnetic resonance response signals of the frequency of the maximum rfsignal and for applying an imaging rf signal to the rf coil at thefrequency of the maximum rf signal propagated from the sweep generator.